1. Field of the Invention
The present invention generally concerns magnetic resonance tomography (MRT) as used in medicine for examination of patients. The present invention more particularly concerns a magnetic resonance tomography apparatus as well as a method for operation of such a magnetic resonance tomography apparatus, with which a high contrast can be achieved for imaging on the basis of a gradient echo sequence.
2. Description of the Prior Art
MRT is based on the physical phenomenon of nuclear magnetic resonance and has been successfully used for over 15 years as an imaging modality in medicine and in biophysics. In this examination modality the subject is exposed to a strong, constant magnetic field. The nuclear spins in the atoms in the subject, which were previously randomly oriented, are aligned.
Radio-frequency energy can now excite these “ordered” nuclear spins to a specific oscillation. This oscillation generates the actual measurement signal in MRT, the measurement signal being acquired by suitable acquisition coils. The measurement subject can be spatially coded in all three spatial directions by the use of non-homogeneous magnetic fields generated by gradient coils. The method allows a free selection of the slice to be imaged, so slice images of the human body can be acquired in all directions. MRT as a slice imaging method in medical diagnostics is distinguished as a non-invasive examination method primarily due to its versatile contrast capability. MRT has developed into a method superior to x-ray computed tomography (CT) due to the exceptional display capability of soft tissue. MRT today is based on the use of spin echo and gradient echo sequences that enable an excellent image quality with measurement times on the order of minutes.
The continuous technical development of the components of MRT apparatuses and the introduction of faster imaging sequences opens MRT to ever more fields of use in medicine. Real-time imaging to support minimally-invasive surgery, functional imaging in neurology and perfusion measurement in cardiology are only a few examples. In spite of the technical progress in the design of MRT apparatuses, image contrast and signal-noise ratio (SNR) of the MRT image remain limiting factors for many applications of MRT in medical diagnostics.
Particularly in the case of image acquisitions of the head, the goal is to ensure a good segmentation, meaning a good contrast between grey brain matter, white brain matter and cerebrospinal fluid (CSF). One possibility for this is a (semi)automated method for segmentation using T1-weighted MPRAGE data. In the following the MPRAGE sequence as well as the associated fundamentals are therefore initially described.
The acquisition of the data in MRT occurs in k-space (frequency domain). The MRT image in the image domain is linked with the MRT data in k-space by a Fourier transformation. The spatial coding of the subject which spans k-space occurs by means of gradients in all three spatial directions. For this purpose, auxiliary magnetic fields Gx, Gy and Gz, whose field strengths linearly depend on the respective spatial coordinates x, y and z, are superimposed on the homogeneous basic magnetic field. Without limitation as to generality, in the further discussion Cartesian k-space is assumed that is scanned (sampled) per slice or per line.
In MRT imaging the gradient fields are used in different ways. In selective slice excitation a gradient field is superimposed on the homogeneous basic field along one of the coordinate axes (typically the z-axis) during the RF pulse. By selection of a specific frequency spectrum of the RF pulse, only nuclei within a specific slice perpendicular to the z-axis are excited. For frequency coding, a gradient field (typically along the x-axis) is superimposed on the basic magnetic field during the acquisition of the RF signal. The readout of the RF signal ensues in N-equidistant time steps Δt. For phase coding a gradient field (typically along the y-axis) with a constant gradient strength is superimposed on the basic magnetic field for a specific time ty before the acquisition of the RF signal. The readout ensues by repetition of the sequence N times, with the gradient strength being increased in equidistant steps per repetition.
FIG. 2A shows the principle of MRT imaging with the 2D Fourier method. A slice is established by selective slice excitation along the z-axis; this slice of k-space is subsequently scanned line-by-line. The data of a single k-space line are frequency-coded by means of a gradient Gy upon readout. The readout of a line ensues in N equidistant time steps. Each line in k-space has the interval Δkx that is generated by a phase coding step. The imaging sequence is repeated N times for various values of the phase coding gradient Gx. In total, a number matrix with N×N data points is obtained, from which an MRT in the image domain can be constructed by 2D Fourier transformation. FIG. 2B shows the 3D Fourier method. The slice-selection gradient is replaced by a second phase coding gradient. This means that the entire volume of the nuclei is excited by the RF pulse and the spatial information is coded exclusively by orthogonal gradients, namely by two phase coding gradients and one frequency coding gradient. M slices perpendicular to the z-axis are acquired, each slice being scanned line-by-line in k-space. The coding within a slice of k-space ensues by a frequency coding gradient in the y-direction as well as a phase coding gradient in the x-direction. A number matrix with M×N×N data packets is thus obtained in total.
FIG. 3 schematically shows the excitation and gradient scheme of the known FLASH (Fast Angle Low Shot) sequence. This is based on the principle of the gradient echo technique. Fast image sequences that are based on the principle of small angle excitation, and in which the echo signals generated exclusively by gradient reversal, are designated as gradient echo sequences (GE sequences). In small angle excitation, flip angles of á<90° are used, but only a small fraction of the longitudinal magnetization is rotated in the transversal plane. Thus it is not necessary to wait as long for the relaxation of the magnetization, which leads to significant time savings. Furthermore, the dephasing of the transverse magnetization caused by the two gradients is compensated by the polarity reversal, such that a gradient echo arises. In FIG. 3 the RF pulse with a small angle excitation below an angle is shown in the first line and the RF signal with the gradient echo is subsequently shown on the time axis. In the second line the slice-selection gradient Gz is plotted along the time. As already explained, the slice-selection gradient is superimposed on the homogeneous magnetic field along the z-axis during the RF pulse and the slice-selection gradient is subsequently reversed in terms of polarity for the purpose of dephasing. In the third line the frequency coding gradient Gy is shown along the time axis. A gradient field in the y-direction is superimposed on the homogeneous magnetic field for the frequency coding after polarity reversal of the gradient during the acquisition of the RF signal. The phase coding gradient Gx is shown along the time axis in line 4. For phase coding along the x-axis, a constant gradient is hereby switched on for a defined time before acquisition of the RF signal and the sequence is repeated Nx times. The transverse magnetization is destroyed after the data acquisition via spoiler gradients switched in each of the three spatial axes after acquisition of the RE signal. The echo time TE designated in FIG. 3 is the time from the radiation of the RF pulse up to the gradient echo and the repetition time TR is the time for a sequence pass.
The scheme of the k-space scan of a gradient echo sequence is shown in FIG. 4. After the RF pulse the signal is located in the center of k-space (1). A dephasing of the signal at the point (2) occurs due to the phase coding gradients and the dephasing in the readout direction, A k-space line is scanned (3, 4) during the reverse-polarized readout gradients and the signal is acquired. The gradient echo occurs at the point (3). The entire process is repeated Nx times with phase coding gradients of various strengths such that an image of the entirety of k-space is generated.
FIG. 5 shows the basic principle of the MPRAGE (Magnetization Prepared Rapid Gradient Echo) sequence. This is based on the 3D Fourier method as well as the magnetization preparation. A preparation phase is activated before the actual image phase to achieve shorter measurement times and a good tissue contrast. The preparation phase effects a preparation of the magnetization that is dependent on the relaxation times T1 and T2. The magnetization prepared in this manner is spatially coded and scanned using the gradient echo sequence. FIG. 5 schematically shows the workflow of the MPRAGE sequence in which a magnetization preparation initially occurs; in the imaging phase all Fourier lines are subsequently acquired in the x-direction given a constant value kz along the z-axis. A recovery phase follows for a better SNR and thus a better contrast, and the sequence is subsequently repeated for further values of kz.
The MPRAGE sequence is used by default for depiction of T1-weighted images of the head with good contrast of the grey and white brain matters. The images are used both for routine clinical examinations, but also increasingly in recent times for automatic determination/segmentation of volumes of the brain, specific brain regions or specific tissue types (morphometry). A requirement for this purpose is that the data produce a good contrast between white brain matter and grey brain matter and a good contrast between cerebrospinal fluid (CSF) and grey brain matter.
In (semi-)automated methods for segmentation using T1-weighted MPRAGE data, the inversion time TI (i.e. the time from the beginning of the sequence up to reaching the k-space center during the scan of the gradient echo sequence) is typically set such that a compromise between grey-white contrast and grey-CSF contrast is made. In case of doubt, a manual segmentation or description of the contrast limits is then required. Moreover, the consequence is that an error in the description of the contrast limits is tolerated. This is a problem particularly at high field strengths, since the B1 homogeneity is generally poorer and makes it difficult to achieve completely automatic segmentation, in addition to causing dielectric resonance effects.